Description

Field of the Invention

[0001]

The invention relates to the field of treatment for cardiac
arrhythmias, and more particularly to methods for delivering the
optimum duration of an electrical shock to the heart during
treatment of cardiac arrhythmias.

Background of the Invention

[0002]

Ventricular fibrillation, an often fatal heart arrhythmia,
can be terminated by the application of one or more electrical
current pulses delivered to the heart through electrodes applied
to the chest or implanted within the body. Since the first use
on humans of a completely implantable cardiac defibrillator in
1980, research has focused on making continually smaller and more
efficient defibrillation devices. In addition, reducing the
defibrillation threshold (DFT) energy level applied to the heart
by the defibrillaton pulses reduces the likelihood of damaging
tissue adjacent to the electrodes.

[0003]

A conventional implantable defibrillator includes an
electrical pulse generator and an arrhythmia detection circuit
coupled to the hear by a series of two or more electrodes
implanted in the body. A battery power supply, and one or more
charge storage capacitors are used for delivering defibrillation
shocks in the form of electrical current pulses to the heart.

[0004]

Currently, the primary constraint in reducing the size of
an implantable defibrillator is reducing the battery size and the
size of the storage capacitor(s). Accordingly, improvements in
the area of implantable defibrillators have focused in two areas:
(1) more efficient defibrillation waveforms, and (2) more
efficient electrode configurations and placements. Stated in
other words, the primary variables that can be adjusted in the
design to lower the shock strength required for defibrillation
include those variables relating to the defibrillation waveform,
such as duration, polarity, and waveshape, and those variables
relating to the electrodes, such as materials, size, shape and
location.

[0005]

An example of a development in the area of electrodes is
U.S. Patent No. 4,827,932 to Ideker et al. which relates to a
pair of spaced apart epicardial implantable defibrillation patch
electrodes. A respective patch electrode is attached over each
of the right and left ventricles in an attempt to achieve a
uniform voltage gradient throughout the entire ventricular mass.

[0006]

In the area of defibrillation waveforms, U.S. Patent No.
4,641,656 to Smits discloses a method of applying a sequence of
defibrillating pulses to the heart from a series of four
electrodes. Two adjacent electrodes have positive polarity and
the other two electrodes have negative polarity in an attempt to
concentrate defibrillation energy in the heart wall rather than
through the centre of the heart. Two or more such pulses are
applied, with a reverse in polarity of one pair of opposing
electrodes between each pulse. Another pulsing scheme is
disclosed wherein the polarity of the four electrodes alternates
with each adjacent electrode, and with all four electrodes used
simultaneously to defibrillate the heart.

[0007]

Other examples of defibrillating waveforms are disclosed in
U.S. Patents No.s 4,637,397 to Jones et al., 4,800,883 to
Winstrom, and 4,821,723 to Baker Jr. et al. These patents
disclose multiphasic defibrillation waveforms wherein the
polarity of pulses is reversed. U.S. Patent No. 4,768,512 to
Imutran relates to a high frequency truncated exponential
waveform. U.S. Patent No. 4,727,877 to Kallok discloses a
transvenous lead configuration wherein a first electrical pulse
is delivered to a first pair of electrodes between the right
ventricular apex and the superior ven cava, and after a
predetermined delay, a second pulse is delivered to a second pair
of electrodes between the right ventricular apex and the coronary
sinus.

[0008]

A further example of a device for delivering a
defibrillating shock is disclosed in European Patent Application
EP-A-0528751. This patent discloses a method of automatically
adjusting the pulse duration of a fixed pulse width truncated
exponential waveform defibrillation shock. However, this is
determined using the impedance measured or calculated following
a delivered shock.

[0009]

None of these efforts, however, sufficiently control the
waveform to maximise the efficiency of the defibrillation pulses
and thereby reduce the risk of damage to adjacent tissue and
minimise the size of batteries, capacitors and other
defibrillator hardware.

Summary of the Invention

[0010]

It is therefore, one object of the present invention to
provide a method for producing an optimum waveform for treating
cardiac arrhythmias of a subject. It is a further object that
the method provide an optimum monophasic or biphasic waveform for
defibrillating the heart of a subject.

[0011]

A further object of the present invention is to provide a
waveform which minimizes damage to the cardiac tissue of the area
of the heart receiving the counter-shock waveform signal.

[0012]

A further aspect of the method comprises delivering a second
truncated waveform signal of opposite polarity to the first
waveform signal to the pair of electrodes.

[0013]

These objects and advantages are achieved by a method of
determining the optimum duration of an electric pulse for
electric counter-shock cardiac arrythmia treatment comprising:

convolving the waveform of an electric pulse counter-shock signal
with the impulse response.of a parallel RC circuit having a model
time constant to approximate the response for a heart, wherein
the model time constant is a value from 1 to 4msec; then

determining the peak response time of the waveform from the
convolved waveform.

Brief Description of the Drawings

[0014]

Figure 1 is a graph of the model response to a monophasic
truncated exponential waveform with varying pulse durations.

Figure 2 is a graph of the model response to a biphasic
truncated exponential waveform with varying second pulse
durations.

Figure 3 is a graph of the model response to an optimal
biphasic truncated exponential waveform according to the present
invention.

Figure 4 is a schematic diagram illustrating a cardiac
defibrillator employing the method according to the present invention.

Detailed Description of the Invention

[0015]

The present invention will now be described more fully
hereinafter with reference to the accompanying drawings, in which
preferred embodiments of the invention are shown.

[0016]

This invention, however, may be embodied in many different
forms and should not be construed as limited to the embodiments
set forth herein. Rather, applicants provide these embodiments
so that this disclosure will be thorough and complete and will
fully convey the scope of the invention as defined in the claims to those of skill in this
art.

[0017]

The method as described below are described with reference
to a cardiac defibrillator. However, as will be appreciated by
one of skill in this art, these methods may be readily extended
to. any apparatus which treats arrhythmias with electric counter-shock.
For example, the principles of the present invention
apply equally well to the treatment of atrial tachyarrhythmias,
atrial fibrillation, reentrant tachyarrhythmias, ventricular
tachycardia, as well as ventricular fibrillation. These
principles aid in any device which delivers
an electric pulse to the heart to treat a cardiac arrythmia.

[0018]

The membrane of the cardiac myocyte serves as both a
resistor and a capacitor to current flow. Cardiac activation
and, perhaps, defibrillation, is thought to occur when an
electrical stimulus changes the transmembrane potential by at
least a certain minimum amount to raise the transmembrane
potential above a particular threshold value. This threshold
transmembrane potential may need only be reached for an
infinitesimal small interval of time, however, the principles of
the present invention can easily be extended if the threshold
potential must be exceeded for a certain minimum period of time
to cause stimulation or defibrillation. The description provided
below is for providing a waveform which reaches the desired
threshold voltage and then immediately begins to return to the
baseline voltage. The man skilled in the art will appreciate
that principles of the present invention may be readily extended
to provide a waveform signal which maintains a voltage above the
desired threshold voltage for a selected finite time by
increasing the duration of the stimulating pulse.

[0019]

The resistance and capacitance of the membrane of the
cardiac myocyte form an RC network. This means that an applied
voltage across the membrane does not appear immediately when the
voltage is applied, but increases as one minus an exponential
function as the capacitor is charged. Once this time constant
is known, the time after the start of the shock at which the
change in transmembrane potential reaches a peak value can be
calculated for any shock waveform. This technique can be used
to determine the waveform that can stimulate and/or defibrillate
with the smallest peak voltage or total energy. As an
alternative to creating the optimum waveform, this technique can
be used to produce the best waveform from those that can be
practically delivered by a defibrillator. For example, internal
defibrillators deliver a truncated exponential waveform by
charging a capacitor within a defibrillator to a predetermined
voltage and then discharging this capacitor through the
defibrillator electrodes for a certain amount of time. the
capacitor is then disconnected from the defibrillation
electrodes, truncating the exponentially decaying shock waveform.

[0020]

Excitable media in general and in particular the heart has
traditionally been modelled as a parallel RC circuit. If a
square wave is used as input to the model, the resulting
relationship is the Lapique equation for stimulation.
V(t)=atm/cm(1-et/tm)
where V(t) is the voltage across the model at the time t, a is
the amplitude of the input square wave in amps, tm is the model
time constant, and cm is the value of the capacitor in the model.

[0021]

This model may be extended to predict V(t) when any shape
waveform are used as input. The threshold duration relationship
for an input waveform may be derived by convolving the impulse
response of a parallel RC circuit with the input waveform. The
derivative of the result of the convolution is then taken and set
to 0 to solve for the peak response time.

[0022]

The above procedure was utilised for truncated exponential
waveforms are used as input. The equation for the response of
a parallel RC circuit to a decaying truncated exponential
waveform is:
V(t) = atstm (et/ts-et/tm)cm (ts-tm)
here ts is the waveform time constant.

[0023]

This equation can be applied to defibrillation. If all the
tissue that is between the two electrodes is represented as a
parallel RC circuit, then a defibrillation threshold can be
defined as that input to the model that raises the model response
V(t) to some fixed response.

[0024]

While Cm is unknown it is only a scaling factor and,
therefore, does not impact on the remainder of the analysis. In
addition, tm is also unknown. From analysis of the Lapique
equation above, it can be shown that: tm = .693 tc where tc is the
chronaxie time. Based upon the reported values for chronaxie
from the literature from defibrillation of 1 msec to 4 msec, then
a range of tm values of from about 1.5 to about 4.5 msec results.
Other methods of determining tm include transmembrane tissue
studies using a double barrelled microelectrode in a tissue bath.
Tissue studies in animals may be analogised to the human heart.

[0025]

While tm varies from subject to subject, for the purposes
of the present example, tm may be approximated as 3 msec. From
these equations, the optimal capacitor for the exponential
waveform can be determined. To maximise V(t) for a given input,
then ts should be set equal to tm. Furthermore, ts is simply RC,
the impedance between the two electrodes multiplied by the
capacitance value of the waveform. Human impedances vary from
about 20Ω to about 80Ω with a mean of 40Ω. Therefore, for the
purposes of the present example, using the mean of 40Ω results
in the optimal capacitor value for the exponential waveform of
75µf. Thus, the optimum capacitance for a defibrillator may be
selected by measuring the resistance of the implanted electrodes
and then selecting a capacitor which makes the RC time constant
of the defibrillator equal to the model time constant.

[0026]

Alternatively, in an internal defibrillator the capacitance
and shock duration can be calculated to minimise either the
voltage to which the capacitor must be charged or the total
energy to which the capacitor is charged based on the RC time
constant of the cardiac membrane. A mean value of this model
time constant tm can be determined experimentally either by
determining the strength-duration relationship for defibrillation
directly in humans or, by extrapolation to humans from animals.
Furthermore, as by way of background information only tm can be
determined for a particular subject from a strength duration
analysis of the subject where the strength duration curve is
defined as an exponential with tm as the time constant of that
function.

[0027]

Capacitor discharge waveforms are truncated in order to
improve defibrillation efficacy. Analysis of the above equations
show that the model response reaches a maximum response at some
time t and then slowly decays back towards zero. If the
truncated time is longer than this time t, then the energy
delivered after time t is wasted since it does not lead to any
increase in V(t). This time t is (log(tm/ts)/(1/ts-a/tm)). Note
that this equation varies time t as a function to ts. The
capacitor in an implantable device is fixed, but the impedance
across the electrodes can vary from patient to patient, and over
time in the same patient. Since patient impedance can be
determined as a shock is being delivered, the optimal pulse
duration can be determined for each shock and the pulse truncated
after that duration. The capacitance of the treatment apparatus
described herein are discussed with reference to a capacitor
value in the device. However, as will be readily appreciated,
the capacitance of the device or apparatus need not come from a
single passive capacitor but may be an effective capacitance of
the device with contributions from a variety of sources.

[0028]

In addition to the above application with respect to
monophasic truncated exponential waveforms, these principles may
be applied to biphasic truncated exponential waveforms as well.
As with the monphasic waveform application, the pulse duration
of a first truncated exponential waveform for a particular
subject is determined by measuring the voltage at the electrodes
and determining ts for the first pulse of a first polarity. The
duration of the pulse is then determined as described above and
the pulse truncated when that duration is reached. In the
biphasic application the second pulse is then applied after the
first pulse. This second pulse is of opposite polarity to the
first pulse and is truncated when the potential of the membrane
is restored to the initial voltage before the first pulse is
applied. This initial voltage may be referred to as a baseline
voltage and may be 0 volts or may be the resting voltage of the
heart. The duration of the second pulse may be determined in a
similar manner as the first pulse. Solving the above equations
for the duration t which results in restoring the potential to
the baseline voltage results in the following equation:
where tm may be the same model time constant as used for the
first pulse or may be unique to the second pulse and ts is
determined for the second pulse from the impedance of the
electrodes and heart and the values of the capacitance for the
second pulse and d1 is the duration of the first pulse.
Preferably the second pulse duration is not longer than 1.5 to
2 times the duration d1 of the first pulse. The second pulse may
then be truncated at the appropriate time to efficiently return
the membrane to the baseline voltage.

[0029]

As seen in Figure 1, the model response to a constant
amplitude monophasic waveform varies as the duration of the pulse
varies. In Figure 1, the amplitude of the response initially
increases with increased pulse duration but ultimately reaches
a maximum and then begins to decay. Thus, in the response
depicted in Figure 1, the energy in the monophasic truncated
exponential waveform after 3 milliseconds does not contribute to
increased peak amplitude and is therefore inefficient.

[0030]

Figure 2 illustrates the model response to a biphasic
waveform having a 6 millisecond first pulse of constant
amplitude. The various dashed lines represent varying durations
of the second pulse of biphasic waveform. The responses depicted
in Figure 2 illustrate that as the duration of the second pulse
increases the energy of the second pulse causes the amplitude of
the response to overshoot the baseline amplitude and is therefore
inefficient.

[0031]

As is seen in Figures 1 and 2, the duration of the first
pulse and the duration of the second pulse dramatically impact
the response waveform. Figure 3 illustrates the optimised
biphasic waveform according to the present invention for various
electrode impedance values. As seen in Figure 3, the peak
response amplitude is achieved for each of the electrode
impedances and the response returns to the baseline value without
any overshoot.

[0032]

While the present invention has been described with respect
to the use of truncated exponential waveforms, the principles of
the present invention apply equally well to other waveforms. The
optimum duration of any waveform may be derived for both the
initial pulse to reach a threshold voltage and for an opposite
polarity pulse to return to the baseline voltage.

[0033]

As shown in the schematic diagram of the figure, the
implantable apparatus 10 incorporating the method of the present
invention includes an electronic circuit 11 contained with an
implantable housing 12. The electronic circuit 11 is connected
by a series of leads 14 and 15 to an electrode configuration 13
including a series of electrodes positioned adjacent portions of
the heart.

[0034]

The electronic circuit 11 includes a conventional ECG
amplified 21 for amplifying sensed cardiac signals. The
amplified cardiac signals are analysed by a conventional
arrhythmia detector 22 which determines if and what type of
arrhythmia is present. The arrhythmia detector 22 may be one of
several types well known to those skilled in the art and is
preferably capable of distinguishing between high rate malignant
tachycardia and ventricular fibrillation so as to deliver lower
energy shocks in the former case than those to be delivered in
the latter case.

[0035]

A capacitor charging circuit 23, in response to a signal
from the arrhythmia detector 22, charges the storage capacitor
24 to a predetermined voltage from the battery 25. The voltage
may be selected prior to implantation of the apparatus 10 or may
be dependent on the determination of the arrhythmia detector 22.
The discharge of the capacitor 24 is controlled by the controller
26, or multi-phasic circuit, such as described in U.S. Patent No.
4,850,357. The capacitor 24 may be a single capacitor or a bank
of parallel connected capacitors of equivalent capacity as would
be readily understood by those skilled in the art.

[0036]

The controller 26 delivers electrical pulses to the
electrodes through a programmable switch 27. As would be readily
understood by those skilled in the art, the voltage waveform
delivered by the capacitor 24 may be a decaying exponential
waveform. The capacitor charger 23, capacitor 24, battery 25,
controller 26 and programmable switch 27 thus form an electrical
pulse generator for the apparatus 10.

[0037]

Upon the generation of a voltage pulse, voltage detector 32
monitors the voltage at electrodes 40 and 41 through leads 16 and
17. This voltage is provided to the duration calculator 30 for
determination of the time constant ts for the pulse. The time
constant is determined by measuring the voltage at the electrodes
40 and 41 and determining the rate of change of the voltage at
the electrodes 40 and 41. It will be understood by one of skill
in this art that from the rate of change of the voltage at the
electrodes 40 and 41, the time constant ts may be readily
determined. It will be further understood by one of skill in the
art that, while the apparatus illustrated in Figure 4 utilises
voltage detection to determine the time constant ts,
alternatively the current through the electrodes 40 and 41 could
be measured and the time constant ts determined from the current.
One method of measuring the current would be to place a low value
of resistance in series with the electrodes 40 and 41 and measure
the voltage across that resistance. This voltage is then
directly proportional to the current through the electrodes.
Once the time constant ts is determined the duration calculator
30 uses ts combined with the appropriate model time constant from
the time constant storage 31 to determine the appropriate
duration of the pulse.

[0038]

For a monophasic system the duration of the pulse is
determined such that the pulse is terminated when maximum voltage
at the tissue is reached. For a biphasic system the duration for
maximum voltage is determined for the first pulse and the
duration to return the tissue to the baseline voltage is
determined for the second pulse of opposite polarity. The
determination of the appropriate duration of the pulses may be
carried out through the use of the principles described above.

[0039]

The pulse duration information may then be transferred from
the duration calculator 30 to the controller 26 which interrupts
the pulse after the desired duration is reached by controlling
programmable switch 27. As will be understood by one of skill
in the art, the voltage detector 32, the duration calculator 30
and the time constant storage 31 may utilise electronic circuits
known to one of skill in the art for measuring voltages, storing
information and performing mathematical calculations. As will
be further understood by one of skill in this art, the voltage
detector 32, the duration calculator 30 and the time constant
storage 31 may comprise a plurality of integrated circuits or may
be incorporated into a single integrated circuit or may be
incorporated into the electronic circuitry of the electric pulse
generator described above.

[0040]

With respect to the time constant storage 31, the values of
the model time constant tm may be incorporated into the device
at the manufacture or may be programmed into the device at the
time of implantation. Thus, the model time constant may be
adjusted for each individual subject. Furthermore, the model
time constant may be updated over time after implantation in a
particular subject. The ability to update the model time
constant after implantation would allow the device to compensate
or adjust for variations over time of the subject's cardiac
membrane RC time constant. Furthermore, one model time constant
may be used for both monophasic and biphasic systems and for both
pulses of biphasic systems or individual model time constants may
be utilised for the first and second pulses of a biphasic system.
In the present invention model time constants have a value of from 1 to
4 milliseconds, preferably from about 2.5 to about
3.5 milliseconds .

[0041]

A further aspect of the present invention involves placing
electrodes 40 and 41 can be placed in contact with the cardiac
membrane of the subject. As will be appreciated by one of skill
in this art, the number and location of the electrodes may be
selected to increase the efficiency of the defibrillator. One
such placement of the electrode is described in U.S. Patent No.
5,224,476 by way of background information only. After placing
the electrodes 40 and 41 and the electrode leads 14 and 15 in the
subject, the RC model time tm constant can be determined for the
subject and the electrode placement by determining the strength-duration
relationship for defibrillation. Various external
apparatus known to one of skill in this art may be connected to
leads 14 and 15 for determining the model time constant. After
determining the time constant it could then be stored in the
implantable electric circuitry 11 in the time constant storage
31. Leads 14 and 15 would then be connected to the electronics
11 and the device may then be implanted in the subject.

[0042]

The cardiac arrythmia apparatus described above has been
described with respect to monophasic and biphasic truncated
exponential waveforms. However, the benefits of the present
invention may also be realised using other waveforms.
Furthermore, the description above relates to the production of
a single shock pulse, however, as will be appreciated by one of
skill in the art, the principles and methods described herein are
equally applicable to multiple pulses or a series of pulses of
varying or similar waveforms.

[0043]

As will be understood by one of skill in this art, the
benefits of the present invention may also be realised in a non-implantable,
external defibrillator. Such a device would have
essentially the same schematic representation as that shown in
Figure 4 but would not be restricted by the size limitations of
the implantable device.

[0044]

The present invention is useful in a method of
defibrillating the heart of a subject. The method involves
providing a truncated exponential waveform to a set of electrodes
positioned so as to defibrillate the heart of a subject. The
voltage at the electrodes is measured during the application of
a truncated exponential waveform signal and a time constant ts is
calculated from the rate of change of the voltage at the
electrodes. This time constant can then be used in conjunction
with a model time constant tm derived for human cardiac membrane
to determine the duration of the pulse which will achieve the
desired peak membrane voltage. The duration of the pulse which
achieves peak membrane voltage may be obtained by solving the
equations described above. The waveform is then interrupted
after a pulse of the desired duration has been delivered to the
electrodes. The apparatus of Figure 4 is of particular benefit
in carrying out the method of the present invention.

[0045]

In addition to the method described above, a second
truncated exponential waveform signal can be applied to the
electrodes to produce a biphasic defibrillating waveform. This
second truncated exponential waveform signal is of opposite
polarity to the first truncated exponential waveform signal.
Optionally, in this alternative method, a second time constant
ts is calculated by measuring the rate of voltage change at the
electrodes during application of the second truncated exponential
waveform signal. Using the second time constant ts and the model
time constant tm the duration of the truncated exponential
waveform which returns the membrane voltage to the baseline
voltage may be determined. The second truncated exponential
waveform is then interrupted when this duration is achieved.
Optionally a second model time constant tm for the second
exponential waveform may be utilised in calculating the desired
duration of the second waveform.

[0046]

In addition to the methods just described, the model time
constant utilised in both the methods employing monophasic
defibrillators and the methods employing biphasic defibrillators
may be adjusted based upon the model waveform response of the
cardiac membrane of a particular subject. The adjustment to the
model time constant may, in the case of an implantable
defibrillator, occur at the time the defibrillator is implanted
in the subject, or, in the case of an external defibrillator at
the time the defibrillator is utilised. Furthermore, in the case
of an implanted defibrillator the model time constant may be
adjusted to compensate for changes in the cardiac membrane of the
subject which occur over time. To determine the model time
constant the following procedures may be utilised either at the
time of implantation, at the time of use or over the life of the
implanted defibrillator.

[0047]

Estimates of tm may be determined from a pacing strength-duration
curve at the time of implantation or periodically by the
device. The pacing strength-duration curve for a truncated
exponential pacing stimuli can then be fit to the following
equation:
Ip = vthcm(ts-tmtstm)e-tts-e-ttm
where tm is the model time constant, cm is the model capacitance,
and ts is the CDW time constant. In the above equation t is the
time when Vth occurs, the shorter of D or In (tm/ts)/(1/ts-1/tm)
(optimal time). Standard curve fitting techniques can be used
to determine tm.

[0048]

It is possible to
select a cardiac arrhythmia treatment apparatus,
such as a cardiac defibrillator, from a set of such apparatus for
implantation in a subject. The selection method includes
providing a set of implantable cardiac arrythmia treatment
apparatus, each of which has a different value for the storage
capacitor which delivers a waveform signal, which may be, for
example, a truncated exponential waveform signal. During
implantation of the cardiac defibrillators a set of electrodes
is positioned to provide an electric counter-shock, such as that
in defibrillation, to the heart of the subject in which the
apparatus is to be implanted. The impedance between the
electrodes is then measured. This impedance is then combined
with the capacitor values of the cardiac arrythmia treatment
apparatus to select the apparatus whose capacitance values
produces an RC time constant which most efficiently produces the
desired waveform in the subject. Preferably, the RC time
constant of the apparatus and the electrodes is equal to the
model time constant for cardiac membrane. More preferably, the
RC time constant is equal to the model time constant for the
cardiac membrane of the subject in which the apparatus is to be
implanted. The model time constant of the subject may be
determined by the methods described above. The selected
apparatus is then implanted in the subject.

Example

[0049]

In the past a capacitor discharge waveform (CDW) has been
approximated by square wave of the same duration (D) and with
amplitude equal to the average current (Iave) of the CDW. In this
approximation, as the CDW gets longer, Iave should approach a
constant. Defibrillation studies have show that for short time
constant CDW's, peak current and not Iave becomes constant as the
CDW becomes longer. Thus, as D in the denominator of the Iave
calculation increase, Iave continues to decrease and does not
approach a constant. Using the above equations the threshold-duration
curves for CDW's with ts = 4, 7, 10 and 15 milliseconds
in six dogs. D was 2, 4, 6, 8 and 10 milliseconds. The data was
fit to the standard strength-duration relationship: Iave =
rheo/(1-e-.693D/t), and to the model relationship discussed above.
Chronaxie was 2.01±0.38 ms (95% confidence interval). Rheobase
was 5.99±0.53A. The correlation coefficient for the strength-duration
relationship was 0.90. The time constant tm was
2.64±0.38 ms. Vth/cm was 14.09±1.23 A. The correlation
coefficient for our model was 0.97. This model predicts that the
optimal D for a CDW is given by the optimal time. For durations
longer than D, V no longer increases and so any energy delivered
after this time does not contribute to defibrillation.

[0050]

The foregoing is illustrative of the present invention as defined by the following claims, and
is not to be construed as limiting thereof.

Claims (2)

A method of determining the optimum duration of an electric
pulse for electric counter-shock cardiac arrythmia treatment
comprising:

convolving the waveform of an electric pulse counter-shock
signal with the impulse response of a parallel RC
circuit having a model time constant to approximate the
response for a heart, wherein the model time constant is a
value from 1 to 4msec; then

determining the peak response time of the waveform from
the convolved waveform.

A method according to claim 1 further comprising the step
of generating an electric counter-shock pulse having a
duration of the determined peak response time.

EP199701020351993-10-061994-09-30Method of determining the optimum duration of an electric pulse for counter-shock treatment of cardiac arrhythmia
Expired - LifetimeEP0782870B1
(en)